TREATMENT MACHINES FOR EXTERNAL BEAM RADIOTHERAPY
5.2. X RAY BEAMS AND X RAY UNITS
Clinical X ray beams typically range in energy between 10 kVp and 50 MV and are produced when electrons with kinetic energies between 10 keV and 50 MeV are decelerated in special metallic targets.
Most of the electron’s kinetic energy is transformed in the target into heat, and a small fraction of the energy is emitted in the form of X ray photons, which are divided into two groups: characteristic X rays and bremsstrahlung X rays.
5.2.1. Characteristic X rays
Characteristic X rays result from Coulomb interactions between the incident electrons and atomic orbital electrons of the target material (collision loss).
In a given Coulomb interaction between the incident electron and an orbital electron, the orbital electron is ejected from its shell and an electron from a higher level shell fills the resulting orbital vacancy. The energy difference between the two shells may either be emitted from the atom in the form of a characteristic photon (characteristic X ray) or transferred to an orbital electron that is ejected from the atom as an Auger electron.
●The fluorescent yield w gives the number of fluorescent (characteristic) photons emitted per vacancy in a shell (0 _< w _< 1) and ranges from zero for low Z atoms through 0.5 for copper (Z = 29) to 0.96 for high Z atoms for K shell vacancies, which are the most prominent sources of character- istic X rays.
●The photons emitted through electronic shell transitions have discrete energies that are characteristic of the particular target atom in which the transitions have occurred; hence the term characteristic radiation.
5.2.2. Bremsstrahlung (continuous) X rays
Bremsstrahlung X rays result from Coulomb interactions between the incident electron and the nuclei of the target material. During the Coulomb interaction between the incident electron and the nucleus, the incident electron is decelerated and loses part of its kinetic energy in the form of bremsstrahlung photons (radiative loss).
●Photons with energies ranging from zero to the kinetic energy of the incident electron may be produced, resulting in a continuous brems- strahlung spectrum;
●The bremsstrahlung spectrum produced in a given X ray target depends on the kinetic energy of the incident electron as well as on the thickness and atomic number Z of the target.
5.2.3. X ray targets
According to the range R of electrons of a given kinetic energy EK in the target material, targets are divided into two main groups: thin and thick.
A thin target has a thickness much smaller than R, while the thickness of a thick target is of the order of R. For thin target radiation, the energy radiated is proportional to the product EKZ, where Z is the atomic number of the target.
The intensity versus photon energy (photon spectrum) is constant from zero to the kinetic energy EK of the incident electron, and zero for all energies above EK.
A thick target may be considered as consisting of a large number of superimposed thin targets. The intensity I(hn) of a thick target spectrum is expressed as:
I(hn) = CZ(EK – hn) (5.1)
where
C is a proportionality constant;
hn is the photon energy.
X rays are used in diagnostic radiology for diagnosis of disease and in radiation oncology (radiotherapy) for treatment of disease. X rays produced by electrons with kinetic energies between 10 keV and 100 keV are called superficial X rays, those with electron kinetic energies between 100 keV and 500 keV are called orthovoltage X rays, while those with electron kinetic energies above 1 MeV are called megavoltage X rays.
Superficial and orthovoltage X rays are produced with X ray tubes (machines), while megavoltage X rays are most commonly produced with linacs and sometimes with betatrons and microtrons.
Typical thin and thick target bremsstrahlung spectra originating from 100 keV electrons striking a thin and thick target, respectively, are shown in Fig. 5.1.
5.2.4. Clinical X ray beams
A typical spectrum of a clinical X ray beam consists of line spectra that are characteristic of the target material and that are superimposed on to the continuous bremsstrahlung spectrum. The bremsstrahlung spectrum originates in the X ray target, while the characteristic line spectra originate in the target and in any attenuators placed into the beam.
●The relative proportion of the number of characteristic photons to bremsstrahlung photons in an X ray beam spectrum varies with the electron beam kinetic energy and atomic number of the target. For example, X ray beams produced in a tungsten target by 100 keV electrons contain about 20% characteristic photons and 80% bremsstrahlung photons, while in the megavoltage range the contribution of characteristic photons to the total spectrum is negligible.
FIG. 5.1. Typical thin target (curve 1) and thick target (curves 2, 3 and 4) spectra for an X ray tube in which 100 keV electrons strike the target. Curve (1) is for a thin target producing a constant intensity for photon energies from zero to the kinetic energy of electrons striking the target (100 keV). Curve (2) represents an unfiltered spectrum (inside the X ray tube) for a thick target and a superposition of numerous thin target spectra; the spectrum of curve (3) is for a beam filtered by an X ray tube window (low energy photons are filtered out); the spectrum of curve (4) is for a beam filtered by the X ray tube window and additional filtration.
●In the diagnostic energy range (10–150 kV) most photons are produced at 90º from the direction of electron acceleration, while in the megavoltage energy range (1–50 MV) most photons are produced in the direction of electron acceleration (forward direction: 0º).
5.2.5. X ray beam quality specifiers
Various parameters, such as photon spectrum, half-value layer (HVL), nominal accelerating potential (NAP) and beam penetration into tissue equivalent media, are used as X ray beam quality indices (see Sections 9.8.1 and 9.8.2 for details):
●A complete X ray spectrum is very difficult to measure; however, it gives the most rigorous description of beam quality.
●The HVL is practical for beam quality description in the superficial (HVL in aluminium) and orthovoltage (HVL in copper) X ray energy range, but not practical in the megavoltage energy range because in this energy range the attenuation coefficient is only a slowly varying function of beam energy.
●The effective energy of a heterogeneous X ray beam is defined as that energy of a monoenergetic photon beam that yields the same HVL as does the heterogeneous beam.
●The NAP is sometimes used for describing the megavoltage beam quality.
The NAP is determined by measuring the ionization ratio in a water phantom at depths of 10 and 20 cm for a 10 × 10 cm2 field at the nominal source to axis distance (SAD) of 100 cm.
●Recent dosimetry protocols recommend the use of tissue–phantom ratios or percentage depth doses (PDDs) at a depth of 10 cm in a water phantom as an indicator of megavoltage beam effective energy (beam quality index).
5.2.6. X ray machines for radiotherapy
Superficial and orthovoltage X rays used in radiotherapy are produced with X ray machines. The main components of a radiotherapeutic X ray machine are: an X ray tube; a ceiling or floor mount for the X ray tube; a target cooling system; a control console; and an X ray power generator. A schematic diagram of a typical therapy X ray tube is shown in Fig. 5.2.
●The electrons producing the X ray beams in the X ray tube (Coolidge tube) originate in the heated filament (cathode) and are accelerated in a
vacuum towards the target (anode) by an essentially constant potential electrostatic field supplied by the X ray generator.
●The efficiency for X ray production in the superficial and orthovoltage energy range is of the order of 1% or less. Most of the electron kinetic energy deposited in the X ray target (~99%) is transformed into heat and must be dissipated through an efficient target cooling system.
●To maximize the X ray yield in the superficial and orthovoltage energy range the target material should have a high atomic number Z and a high melting point.
●With X ray tubes, the patient dose is delivered using a timer and the treatment time must incorporate the shutter correction time (see Section 6.16), which accounts for the time required for the power supply components to attain the steady state operating conditions.
●The X ray tube current is controlled by a hot filament emission of electrons, which, in turn, is controlled by the filament temperature (thermionic emission). For a given filament temperature the X ray tube current increases with the tube (anode) voltage, first rising linearly with voltage in the space charge limited region and saturating at higher voltages when all electrons emitted from the cathode are pulled to the anode.
FIG. 5.2. Typical therapy X ray tube (reprinted from Johns, H.E., and Cunningham, J.R., with permission).
●Research is currently being carried out on cold field emission cathodes produced with carbon nanotubes (CNTs). The CNT based cold cathode X ray technology may lead to more durable as well as miniature and portable X ray sources for industrial and medical applications.
5.3. GAMMA RAY BEAMS AND GAMMA RAY UNITS